It has become increasingly common to treat a variety of medical conditions by introducing a medical device into the vascular system or other lumen within a human or veterinary patient such as the esophagus, trachea, colon, biliary tract, or urinary tract. For example, medical devices used for the treatment of vascular disease include stents, catheters, balloon catheters, guide wires, cannulas and the like. While these medical devices initially appear successful, the benefits are often compromised by the occurrence of complications, such as late thrombosis, or recurrence of disease, such as stenosis (restenosis), after such treatment.
Restenosis, for example, involves a physiological response to the vascular injury caused by angioplasty. Over time, de-endotheliazation and injury to smooth muscle cells results in thrombus deposition, leukocyte and macrophage infiltration, smooth muscle cell proliferation/migration, fibrosis and extracellular matrix deposition. Inflammation plays a pivotal role linking early vascular injury to the eventual consequence of neointimal growth and lumen compromise. In balloon-injured arteries, leukocyte recruitment is confined to early neutrophil infiltration, while in stented arteries, early neutrophil recruitment is followed by prolonged macrophage accumulation. The widespread use of coronary stents has altered the vascular response to injury by causing a more intense and prolonged inflammatory state, due to chronic irritation from the implanted foreign body, and in the case of drug eluting stents (DES), from insufficient biocompatibility of the polymer coating.
Over the past several years, numerous local drug delivery systems have been developed for the treatment and/or the prevention of restenosis after balloon angioplasty or stenting. Examples include local drug delivery catheters, delivery balloon catheters, and polymeric drug coated stents. Given that many diseases affect a specific local site or organ within the body, it is advantageous to preferentially treat only the affected area. This avoids high systemic drug levels, which may result in adverse side effects, and concentrates therapeutic agents in the local area where they are needed. By treating just the diseased tissue, the total quantity of drug used may be significantly reduced. Moreover, local drug delivery may allow for the use of certain effective therapeutic agents, which have previously been considered too toxic or non-specific to use systemically.
One example of a local delivery system is a drug eluting stent (DES). The stent is coated with a polymer into which drug is impregnated. When the stent is inserted into a blood vessel, the polymer degrades and the drug is slowly released. The slow release of the drug, which takes place over a period of weeks to months, has been reported as one of the main advantages of using DES. However, while slow release may be advantageous in the case where a foreign body, such as a stent, is deployed, which is a source of chronic irritation and inflammation, if a foreign body is not implanted it is instead advantageous to rapidly deliver drug to the vascular tissue at the time of treatment to inhibit inflammation and cellular proliferation following acute injury. Thus, a considerable disadvantage of a DES, or any other implanted medical device designed for sustained release of a drug, is that the drug is incapable of being rapidly released into the vessel.
Additionally, while drug-eluting stents were initially shown to be an effective technique for reducing and preventing restenosis, recently their efficacy and safety have been questioned. A life-threatening complication of the technology, late thrombosis, has emerged as a major concern. Drug eluting stents cause substantial impairment of arterial healing, characterized by a lack of complete re-endothelialization and a persistence of fibrin when compared to bare metal stents (BMS), which is understood to be the underlying cause of late DES thrombosis. Concerns have also been raised that the polymeric matrix on the stent in which the anti-proliferative drug is embedded might exacerbate inflammation and thrombosis, since the polymers used are not sufficiently biocompatible. These polymeric systems are designed to facilitate long-term sustained release of drug over a period of days, months, or years, not over a period of seconds or minutes. These polymeric drug coatings of medical devices do not release the polymer, which remains on the device even after drug is released. Even if biodegradable polymers are used, polymer and drug are not released at the same time. Rapid release of drug, an intent of embodiments of the present invention, from these polymeric systems is not possible. Thus, combining a therapeutic agent with a polymer in a medical device coating may have significant disadvantages.
Another important limitation of the DES is that the water insoluble drugs are not evenly distributed in the polymeric matrix of the coating. Furthermore, drug and polymer are concentrated on the struts of the stent, but not in gaps between the struts. The non-uniform distribution of drug causes non-uniform drug release to the tissue of the vessel walls. This may cause tissue damage and thrombosis in areas exposed to excess drug and hyperplasia and restenosis areas that are undertreated. Thus, there is a need to improve the uniformity of drug delivery to target tissues by improving drug solubility in coatings of medical devices by increasing the drug's compatibility with carriers in the coatings, such as a polymeric matrix, thereby eliminating or reducing the size of drug crystal particles in the polymeric matrix or other coating to create a uniform drug distribution in the drug coating on the medical device.
Yet another important limitation of the DES is that only a limited amount of an active agent can be loaded into the relatively small surface area of the stent.
Non-stent based local delivery systems, such as balloon catheters, have also been effective in the treatment and prevention of restenosis. The balloon is coated with an active agent, and when the blood vessel is dilated, the balloon is pressed against the vessel wall to deliver the active agent. Thus, when balloon catheters are used, it is advantageous for the drug in the coating to be rapidly released and absorbed by blood vessel tissues. Any component of the coating that inhibits rapid release, such as a crosslinked polymer, liposome, or encapsulating particle, is necessarily disadvantageous to the intended use of the balloon catheter, which is inflated for a very brief period of time and then removed from the body.
Hydrophilic drugs, such as heparin, have been reported to be deliverable by polymeric hydrogel coated balloon catheters. However, a polymeric hydrogel coating can not effectively deliver water insoluble drugs (such as paclitaxel and rapamycin), because they can not mix with the hydrogel coating. Furthermore, the cross-linked polymeric hydrogel remains on the balloon after drug is released
Alternatively, balloon catheters are reported to have been coated with hydrophobic therapeutic agents that have been encapsulated in particles such as micelles, liposomes, nanoparticles or polymers. Liposomes, which usually contain a core of aqueous solution, and micelles, which do not, have both been reported to be useful for pharmaceutical preparations of water-insoluble drugs for venous injection. However, for purposes of a medical device coating, all of these drug delivery formulations have significant disadvantages: drug loading is poor, and drug release from these preparations is slow.
Oils and lipids mix well with water-insoluble drugs such as paclitaxel or rapamycin, but when micelles or liposomes are then formed by interaction with aqueous media, the particles and particle sizes are relatively unstable, ranging in a broad particle size distribution from several hundred nanometers to several microns in diameter. Several reports suggest that the maximal concentration ratio of drug to lipid that can be stably achieved in these particles is in the range of 0.2 to 0.3; it is often less than 0.1.
Another disadvantage of oil-based liposome formulations is the dependence of drug absorption on the rate and extent of lipolysis. Lipolysis of oil-based triglycerides is difficult and dependent upon many factors, and triglycerides must be digested and drug released in order to be absorbed by diseased tissue. The amount of hydrophobic drug delivered to tissues by these agents will be low, because liposomes and micelles cannot efficiently release hydrophobic drug, which they carry away before it can be absorbed by tissues. Micelles, liposomes, or particles of oils and lipids are therefore not effective at rapidly and efficiently facilitating tissue uptake of drug during a very brief device deployment time, and no report has shown these types of coatings to be effective.
Loading capacity of conventional micelles and liposomes is low. In the absence of other considerations, the highest achievable drug to lipid ratio is advantageous, since high lipid doses may raise concerns of toxicitiy, and it is the drug—not the lipid—that provides the therapeutic benefit, once it is delivered to target tissue. The ratio of drug to lipid in these formulations is often less than 0.1 and almost always less than 0.2-0.3, because a significantly higher concentration of lipid than drug is required in order for the drug to be encapsulated in the particles, miscelles, or liposomes. Formulation stability is highly dependent on drug-lipid interactions that are concentration dependent. “In many studies, a maximum of 3 to 4 mol % drug (with respect to phospholipid) possess stability of sufficient duration as to be clinically useable. 8 mol % paclitaxel liposomes may be physically stable for 15 min or less”-Preparation and Characterization of Taxane-Containing Liposomes. Methods Enzymol. 2005; 391:97-117, p. 101-2. Several attempts to achieve drug to lipid concentration ratios higher than 0.2-0.3 failed (for example, see PDA J Pharm Sci Technol 2006 60(3):144-55). These technologies involve forming the drug/lipid particles first and then coating medical devices with the prepared particles. Liposomes, for example, are prepared by first mixing drug and lipid in organic solvent, then removing the organic solvent to form a lipid film or cake, then hydrating with aqueous solution and sonicating.
There are several reports showing that drug release from liposomal, miscellar, or particular oil/lipid formulations occurs very slowly, in the range of days to weeks or months. In addition, the inventor has found that drug release occurs far too slowly from a coating consisting essentially of just oil/lipid and lipophilic drug, because they bind to each other and to the external surface of the medical device so tightly that the drug cannot rapidly elute off the device during several minutes or less of deployment at the target site. This slow-releasing property of prior approaches to drug:oil, drug:fatty acid, and drug:lipid formulations is not desirable in embodiments of the present inventions, wherein drug release takes place in the range of seconds to minutes. Thus the technology for oil/lipid formulation needs to be improved significantly in order to be useful in the rapid drug release coatings for medical devices of the present invention.
Drug that is encapsulated in polymeric particles may take even longer to diffuse from the coating (the reported range is months to years) and will have further difficulty permeating target tissues rapidly. Microspheres formed with polymeric materials, such as polyesters, when used to encapsulate water insoluble drugs, are unable to release the drug until the polymeric material is degraded. Thus, these polymeric microspheres are useful for sustained release of drug over a long period of time but cannot rapidly release drug and facilitate tissue uptake.
The iodine contrast agent iopromide has been used with paclitaxel to coat balloon catheters and has had some variable success in the treatment of restenosis. Notwithstanding this result, prior approaches to coating formulations with iodine contrast agent and a lipophilic or water-insoluble therapeutic agent have serious drawbacks. The physical characteristics of these coatings are very poor. There is a several-fold variability in drug levels in tissue after treatment, which results in inconsistent therapeutic outcomes.
Experimentally, the present inventor has found that the compatability or miscibility of lipophilic or water-insoluble therapeutic agents and iodine contrast agents is not good, and the integrity and uniformity of the coating is very poor. Particles from the coating easily flake off and are lost during handling. When the coating consists essentially of iopromide and paclitaxel, for example, there is experimentally several-fold variability in drug concentration distribution on the surface of the medical device. Furthermore, there is marked variability in adhesion of drug to the surface of the device, and clumps of contrast agent and drug of variable size are formed that easily flake off the device during minimal handling in the interventional cardiology suite. These deficiencies adversely affect the consistency and uniformity of drug delivery to target tissues, which results in highly variable and inconsistent therapeutic effects and decreased therapeutic utility of the coated device. Improved coatings are therefore needed, coatings that maintain integrity during handling and more effectively and uniformly deliver drug and facilitate its absorption by tissue.
Thus, the technology for formulating coatings of contrast agent and drug needs to be improved significantly in order to achieve consistent therapeutic results and to improve the clinical utility of the rapid drug release coatings for medical devices.
Combining drugs and medical devices is a complicated area of technology. It involves the usual formulation challenges, such as those of oral or injectable pharmaceuticals, together with the added challenge of maintaining drug adherence to the medical device until it reaches the target site and subsequently delivering the drug to the target tissues with the desired release and absorption kinetics. Drug coatings of medical devices must also have properties such that they do not crack upon expansion and contraction of the device, for example, of a balloon catheter or a stent. Furthermore, coatings must not impair functional performance such as burst pressure and compliance of balloons or the radial strength of self- or balloon-expanded stents. The coating thickness must also be kept to a minimum, since a thick coating would increase the medical device's profile and lead to poor trackability and deliverability. These coatings generally contain almost no liquid chemicals, which typically are often used to stabilize drugs. Thus, formulations that are effective with pills or injectables might not work at all with coatings of medical device. If the drug releases from the device too easily, it may be lost during device delivery before it can be deployed at the target site, or it may burst off the device during the initial phase of inflation and wash away before being pressed into contact with target tissue of a body lumen wall. If the drug adheres too strongly, the device may be withdrawn before the drug can be released and absorbed by tissues at the target tissues.
Thus, there is still a need to develop improved coatings of therapeutic agent and contrast agent for medical devices that rapidly, uniformly, and consistently deliver therapeutic agent directly into a localized tissue area during or following a medical procedure, during a brief 0.1 to 2 minute deployment of the device at the target site, so as to treat or prevent vascular and nonvascular diseases such as stenosis. There is still a need to develop coatings with improved compatability of contrast agent and lipophilic or water insoluble drug. There is still a need to develop coatings that preserve coating integrity and uniformity of drug distribution and that diminish the propensity for particles to flake off from the coating during handling. And there is still a need to develop coatings of contrast agent and drug that result in more uniform and consistent delivery of drug to target tissue.
Thus, the technology for formulating medical device coatings of contrast agent and therapeutic agent needs to be improved significantly in order to achieve consistent therapeutic results and to improve the clinical utility of the coated medical devices, which is an object of embodiments of the present invention.